GB2357149A

GB2357149A - MRI using non-homogeneous static field

Info

Publication number: GB2357149A
Application number: GB9929070A
Authority: GB
Inventors: Eres Golan
Original assignee: Topspin Medical Israel Ltd
Current assignee: Topspin Medical Israel Ltd
Priority date: 1999-12-08
Filing date: 1999-12-08
Publication date: 2001-06-13
Also published as: WO2001042807A2; US6600319B2; IL149945A0; EP1244922A2; GB9929070D0; US20010011889A1; JP2003516214A; AU1727101A; WO2001042807A3

Abstract

An NMR probe 3 for imaging external surrounding material produces a non-homogeneous static magnetic field in a region of the material, which is then excited using broad band RF pulses. This allows the excited region 5 to have a greater volume than in conventional homogeneous field arrangements, and so the total received signal is increased. The probe may be integrated into an intravascular catheter and used for imaging blood vessel walls.

Description

2357149 - I - Magnetic Resonance Im2ging Device

FIELD Or, THE INVENTION

This invention is generally in the field of Nuclear Magnetic Resonance (NNM) based techniques, and relates to a device and method for magnetic resonance imagging (MR-T). -AJthough not limited thereto, the invention is particularly usefW far medical purposes, to acquire images of cavities in the human body, but may also be used in any industrial application.

jo BACKGROU10 OF THE'MTNTION MM is a known imaging technique, used especially in cases where soft tissues are to be differentiated. Alternative techniques,, such as ultrasound or X-rky based techniques, which mostly utilize spatial variations in material density, have inherently limited capabilities in differentiatin2 soft tissues.

N-MR is a term, used to describe the physical phenomenon in which nuclei, when placed in a static magriet5c field, respond to a superimposed altemating magnetic field. It is knoAm that when the RF magnetic field has a component directed perpendicular to the static magnetic field, and when this component oscillates at a fi-equency known as the resonance frequency of the nuclei th.en the nuclei can be excited by the RF magne#c field. This excitation is manifested in the temporal behavior of nuclear magnetization following the excitation phase, which in turn can be detected by a reception coil and termed the NIM signal. A key element in the utilization of NUR for imaging purposes is that the resonance frequency. known ai the Larmor frequency, has a linew dependence on the intensity of the static magnedc field in which the nuclei reside. By applying a static magnetic field of which the intensity is spatially dependent it is possible to differentiate signals received from nuclei residing,in diff-erent maimetic field intensities. and

C5 - therefore in different spatial locations. The techniques which utilize NI\R phenomena for obtaining spatial distribution images of nuclei and nuclear characteristics are t=ed NIRI.

in conventional MRI techniques, spatial resolution is achievcd by superimposing a stationary magnetic field gradient on a static homogeneous magnetic field. By using a series of excitations and signal receptions under various gradient orientations, a complete image of nuclear distdbution can be obtained.

3.5 Furthermore, it is a unique. quality of MRI that the spatial distribution of chemical and physical characteristics of materials, such as biological tissue, can be. enhanced and contrasted in many different manners by varying the excitation scheme, . kno-;;,,n as the MFJ sequence, and by using an appropriate processing method.

The commercial applidation of'MRI t-ec ques su&rs from the follo-tAirig two basic. drawbacks: the expenses involved with. purchasing and. operating an MFJ setup; and the relatively low signal sensitivity, which requires long image acquisition time. Both of these drawbacks are linked to the requirement in standard MRI techniques to image relatively large volumes, such as the human body. This necessitates oducing a highly homogeneous magnetic field over the entire imaged pr volume, thereby requiring extensive equipment. Additionally, the unavoidable distance between a signal receiving coil and most of the imaging volume significantly reduces imaging "-.Iitivky.

There are a number of applications in which there is a need for imaging relatively small -volumes, where some of the above-noted shortcomings may be so overcome. One such application is geophysical well logging, where the "Whole body"' NRI approach is obviously impossible. Here, a liole is drilled in the earths crust, and measuring equipment is inserted thereinto for local aging of the surrounding medium at different depths.

Several methods and appiratuses have been developed, aimed at extracting NMR data from. the bor.e hole'walls,. including US Patents Nos. 4,350,955; 4v629,986; 4,717,877, 4,717,878; 4,717,876; 5,21q.,447;.5,280,243; and "Remote Inside Oz&' A'AdR ", J. Map. Res., 4 1, p. 400, 1980; "Novel Alla Apparaw for investigating an External Sample Meiaberg et al.. J. Magn. Res.. 97, p. 466, 1.0 The apparatuses disclosed in the above documents'are based on several permanent magnet configurations designed to create relatively homogeneous static magnetic fields in a re on external to the apparatus itself. RF coils are r gi ypically used in such apparatuses to excite the nuclei in the.homogeneous region and. in turv receive the created NINa signal. To create an external region of a homogeneous m"c field, the magnetic configurations have to be carefully designed to reconcile the faci.thal small deviations in structure may'have a disastrous effect on magnetic beld homogeneity. It turns out that such a region of a A-- homogeneous maguetic field can be created only within a narrow radial distance around a fixed position relative to the magnet configuration, and that the charactcristic magnetic field intensities. crmted in this region are gencrOy lcm,. As a result. such apparatuses, although permitting NMR measurements. 'have only limited use as imaging probes for imaging exl=sive regions. of b ore-hole walls.

With respect to medical MRI-based applications, the potential of using an intra-cavit, receiver coil has been investigated, and is disclosed, for example, in the following publications: Kandarpa et dl.,, J. Vase. and Interventional Radiology, 4. p- 419-427, 1993; and US Patent No. 5,699,801. Different designs fol catheter-based receiver coils are proposed for insertion into body cavities, such as arteries- during' interventional procedures. These coils, when located close to the region of interes improve reception sensitivity, thus allowing high- resolution imaging of these regions. Notwithstanding the far that t approach enables the t his resolution to substntially improved, it still suffers from two major drawbacks:

(1) the need. for bulk external setup in order to create the static homogeneous magnetic field and to trarismit the RF excitation signal; and (2) the need to maintain the orientation of the coil axis within certain li relalive to the external magnetic field, in order to ensure satisfactory image quality. Because of these two limitations, the concept of an intra-cavity receiver coil is only half-way towards designing, a fully autonomous intra-cavity imaging probe.

US Patent No. 5,572,132 discloses a concept of combining, the static magnetic field source with the RF coil in a self contained intra-cavity medical imaging probe. Here, several permanent magnet configurations are proposed for creating a-homogeneous magnetic field region external to the imaging probe itself a manner somewhat analogous to the concept upon which the bore-hole apparatuses are based. Also disclosed in tis patent are several RF and gradient coil configurations that may be integrated in the imaging probe in order to allow autonomous imaging capabilities. The suggested' configurations. nevertheless, suffer from the same problems discussed above with respect to the bore- hole apparatuses, namely: a fixed and narrow, homogeneous region to which imaging is limited, and low magnetic field values characteristic of homogeneous mametic field confi=ations.

SUNEYLALRY OF THE INVENMON Ybere is accordingly a need in the art to improve MRI based techniques, by providing a fWly autonomous intra-cavity MRI probe and an imagin _zzm,s 2, m e th o d.

nie present invention. is based on the realization that rather than attempting to overcome problems of non-homogeneity of the magaetic field, this non-homogeneity mkv be used to the advantage of high-resolution imaging. The imaging), probe according to the invention comprises all components necessary to allow maspetic resonance measurements and imaging of Iocal surrouridings- of the probe, obviating the need for external ma gGetic field sources. T.he imaging method is based on. the non-homogeneous sfatic magnetic field created by permanent

1 -5- magnets and on a high sensitivity fF coil block, all located in the inikizing probe itself. This makes the irnksdng probe an autonomos higli-resolution magnetic resonance imaging device, capable of imaging the medium surrounding the probe.

there is thus provided ac&brding to one aspect of ihe present invention, a method for detecting NMR signals coming from a medium., the method comprising:

(i) producing a primary, substantially non-homogeneous, external magnetic field in the medium; (ii) detecting 1he magnetic resonance signals from -%rithin at least one region of said pi#ary, substantially non-homogeneous magnetic field.

The method enables the simultaneous detection of NNT, sianals originalina from nuclei residing in. the non-homogeneous primary magnetic field, and characterized by a substantially wide frequency range with respect to a. mean frequency value. The term "substantially wide fi-eqwncy range" used herein sigpifes; the -Nide frequency range as compared to that utilized by conventional techniques,. which is limited to 1% of the mean value. Thus, the wide frequency range is, generally, more than I % of the mean value.

According to another aspect of the presmt invention, there is provided a probe for producing Nuclear Magnetic Resonance (NTMR) signals coming from a medium surrounding the probe and detecting the produced signals, the probe comprising:

- a magriedc field-forming assembly that produces a primary, subsbmtially non-homogeneous, external magnetic field; and - a transceiver unit comprising; at least one coil block capable of detecting 2 magnetic resonance signals within at least one region of said prima7 magnetic field, said -it least one region extending from the probe up to distances substantially of the probe dimensions.

The receiving coil block has sufficiently high sensitivity, namely such a sensitivity that enables the signaJ detection with sufficiently high signal-to-noise ratio from even a small volume cell of the medium located at a substantially large distance from the probe. For example, the following condition is indicative of the sufficiently high sensitivity: the NMR signal-to-noise ratio per volume cell, (voxel) accumulated in a time tame of less than one minute (constituting th c averaging time period) exceedmg a value of five can be obtained, wherein the lMR signal part is originated from nuclei in the voxel of O.IxO.2xl= in size located in the primary magnetic field at a distance up to the probe dimensions between the center of the voxel and he closest part of the probe. The noise part relates to the noise level of the coil block integrated over a frequency bandwidth equal to the bandwidth of the NMR signal originating froT the same voxel.

Preferably, the coil block is composed of an RF coil, typically wound around a substantially toroidal core. 'ne at least one region of the sufficiently high sensitivity is provided by forming the coil block with at least one narrow core _Zap, so that the gap plane is aligned substantially parallel to the direction of the primary magnetic field in the region of sufficiently high sensitivity. Several such spaced-apart core gaps may be provided so that more than Ope region of sufficiently ffigh se-nsitivily can be created. The RF coil block may serve for both. signal transmission (i.e., generates an RF magnetic field for exciting the nuclei), and for signal reception. Alternatively, the transceiver may comprise a separate element (e.g. coil block) for ggenerating the t-msmission RF magnetic field.

Preferably, the field-forming assembly comprises two permanent magnets (which may be shaped as cylindrical rings) positioned in an axial. spacedapart relationship on a common cylindrical. fmomkmetic core, with their symmetry axes coinciding and defining the Z-axis. and a small inter-magnet gap rem between the magnets. The magnets are magnetized along the X-axis perpendicular to the Z-axis and in opposite directions to each other, that is the N- and 5-poles of one magnet face, respectively, the. S- aria N-poles of the other magnet The assembly formed by the permanent magnets on the magnetic core creates the primary static magnetic field, which CXtemally to the probe assembly and in the etry plane perpendicular to the Z-axis (hereinafter at times the "X-Y- plane' is directed substantially in the +1-7. direction with a maximum intensity obtained along the X-axis. The operation of the probe defines an imagg or measurement slice as the region where changes in the magnetic field component along the Z-axis are substantially small. with.respect to changes in the position alone the Z-axis. By changing the size of the inter-magnet a, the profile of the static magnetic field in gap the imaging slice can be varied. The RF coil block is preferably located in the gap between the magnets.

As for the RF coil b lk, when used for signal reception, the at least one region of the primary magnetic field from which the magnetic resonance signals are detected is located in proximity of the coil gap(s). In.this region, the coil is substantially sensitive to variations in the transverse nuclear maenetization, Le., in the X-or Y-component,,depending on the RF-coil winding method. If the reception RF-coil block is used for transmission as wel. the magnetic flux lines of. the transmission magnetic field are substantially perpendicular to the Z-axis, and the transmission magnetic field intensity is highest in proximity of the coil gap(s). For both reception and transmission purposes, the at least one region. of sufficient reception sefisitivity and of maximum transmission intensity, respectively, can be .dsua&,ed as at least one angular segment of the X-Y-plane, which is syrnmetrical with respect to the X-axis.

The field-forming assembly a nd the RF coil block mkv have the capabihty of revolving, together or separately. around the Z-wds. This revolution results in rotation of the angular segment(s), thereby scanning the imaging slice.

The device may also comprise g gradient coil to create variable magnetic field gradients in the lateral direction, i.e. perpendicularly to the radial direction.and to the Z- axis (a so-called " -,g gradient coil 'Me static etic field strength decreases with the increase of radial distance from the Z-axis, creating a substantially static magnetic field gradient in the radial direction. This can be utilized for creating a radial image dimension.

When the probe is positioned at a fixed angle, i.e. without rotation about the Z-axis, the device can be used for obtaining pseudo-one-dimensional cross- sections cif the surrounding medium residing subst=tially along te X-axis, i.e., zlong the radial passing through the RF-coil gap. An additional image dimension may be c; eated by several techniques. As indicated above, one such technique utilizes a - _zradicnt coil that can be added to-the probe confimmation, as a means to achieve lateral separation (i.e. in the direction). Alternatively, lateral separation may, be achieved by other methods, such as angular selective excitation or special signal processing.

Preferably, the probe is slowly rplated. around - the symmetry axis. A wide-band (Don-selective)- multiple-spip-echo, excitation scheme may be used to acquire the magnetic. resonance signal. created by nuclei in wide overlapping an.gWar sectors external to the probe. The signals acquired may then be veraged in jo order to improve the signal-to-noise mtio, and processed to create a two-dimensional image in polar coordinates (r..

The probe may be integrated into an intravascular cathuer and used for imaging a series of 2-D cross-sections of blood vessel walls. The cross section images will extend appreciably into the vessel walls, providing high- resolution characterization of wall rqorphology, such as the structure and content of existing atherosclerotic lesions.

The imaging probe of the invention for intravascular medical use is included within a catheter thathas preferably at least one hollow lumen to permit flow of blood theretbrough. Such a hollow lumen may be achieved by the use of a hollow, e.g. tubular,. aaggnetic c= supporting the two magnetic field forming members (e.g., permanent magnets). A catheter probe in accordance with an embodiment, of the invention comprises at least one inflatable balloon. When inflated. the balloon fbxes the imaging probe to the vessel walls. minimizing relative motion during the time of image acquisition, while not obstructing the blood circulation.

71e imaging probe may have a variety of different designs to meet particular applications, all being within the scope of the invention as defined herein.

Variations may be in shape, cross section (tubular, cylindrical, rectarLgular, polygonal, etc.), proportions, size... matenal properties (mechanic4 elec-ftromagnetic, etc.) and the Ue.

For ex=ple, the static field Wic with respect to rotations about the

Z,axis can be created. This can be achieved by magnetizing the "Uppee' magnet (i.e., one'of the two magnets positioned on the Z-axis) radially outward and the ,lowe? magnet radially inward, leaving the ferroetic care unchanged. It can flirther obviate the need for rotating the rnagnetic structure.

Multiple high-intensity RF field sepnents or segments of substantially high sensitivity can be created, directed along a series of angles (forexarnple: along 0, 3 15 degrees) covering the W-plane. This will obviate the need for 45, 90,...:, rotating the RF-coil block.

)01 A "stack-' of RF-coil blocks or of mamets/RF-coil blocks aligned along the Z-axis can' be provided for the simultaneous imaging of multiple slices (X-Y-planes).

The preferred structure in terms of static mag3.etic field strength is the use of the original permanent magnets, having only solid cylindrical shapes (not rings) ].5 and no ferromagnefic core, or with a ferr=aeletc core connecting two periphery regions of the magnets (not centered around the Z-a>ds). In this con- figuration, the RF-coil block. is not necessarily a perfect toroid, but it is again iii the inter-magne gap and the entire magnets-coil combination revolves., since both field pattems are not sy=ede for rotations around the Z-axis.

All the above configurations can be made solid, namely without an internal Iumen for blood flow., When such configurations are used for intravascular applications, blood flow can be allowed externaDy to the probe.

The abovedescribed transceiver unit comprising a coil block having sufficiently high sensitivity for receiving magnetic resonance signals within at least :!5 one region of the primary.magnetic field can be advantageously.used in any MRP or NMR-aimed device,' ective of the homogeneity or ncrn-homogencitv of e Irresp th primary mapietic field. which can be created by an external static magnetic field source.

There is. thus provided according to yet another aspect of the present invention, a transceiver unit for use'in a probe for detecting NMR signals of a surrounding, medium, the transceiver unit comprising at least one coil block capable of detecting magnetic. resonance signals within at least one region of a primary external magnetic fielk wherein."id coil block crympri wound m a substantially toroidal core having at least one core gap.

According to yet another aspect of the present invention, there is provided a device for NNR measurements or magnetic resonance imaging of a mediun-4 the device comprising an imaging probe to be located in the vicinity o said - medium and connected to a control station for generating transmission pulses, and for receiving, processing and displayirig, data generated by the probe, the probe jo comprising:- (a) a magnetic fleld'forming assembly that produces a primary substantially non hornoggeneous, external magnetic field in the medium; and (b) a transceiver unit cdmprising at least one coil capable of detecting magnetic resonance signals from within at least one re&n of said 1.5 primary, substantially non-homogeneous magnetic field, said at least one region- extending from the probe up to distinces, substantial.)y of the probe dimezisions.

Imaging of human blood vessels is a preferred embodiment of the invention and the description bellow refers specifically thereto. It should, however be undoubtedly clear tat the fbHowing description of the preferred embodiments does not limit the present invention, but rather serves only to illustrate the invention. It is clear that by routine design Modifications, which are within the reach of the artisan, probes in accordance with the invention for other applications may be designed.

BRIEF DESCRIPTION OF THE DRAWINGS

In order -to understand the invention and to see how it may be carried out in, practice, a prefured embodiment will now be described, by way of nonlbniting exarnple only, with reference to the accompanying drawings, in which:

Fig. I is a schematic illustration of a device according to the invention utilizing an imaging probe integrated in'an intravascular catheter; Fig. ZA more specifically illustrates the main components of the imaging probe of Fig. 1; Fig. 213 is an enlarged view 6f an RF coil of the imaging probe of Fig, 2A; Fig. 3 is a cross-sectional view of the imaging probe of Fig. 2A showing the static magnetic field lines created by the primary magnetic field source;

Fig. 4 is a cross-sectional view of the RF-coil of Fig. 2B, showing the RF magnetic field lines created when current is driven into the coil;

Fig. 5 exemplifies'the static magnetic field profile calculated along a radial vector for the imaging probe configuratioD of Fig. 2A having an outer diameter. of jo 3mm; Fig. 6 exemplifies the RF map'ietic field profile calculated along a radial vector. for the imaging probe &on:5gL=ion of Fig. 2A, having an outer diameter 6f 3mrn4 Figs. 7 to 9 graphically illustrate the operadonal. principles of the imaging probe 4ccordingto the invention; and Figs. 10a to 10c illustrate three more embodiments of & imkgingg probe, respectively, suitable to be used in the device of Fig. 1.

DETAILED DESCKEPTION OF A PREFERRED EMBODEWENT Referring to Fig. 1, there is illustrated an MRI-based system for intravascular imaging,, generally designated 1, utilizing an intravascular catheter 2 that comprises an imaging probe 3 and Mo infiatable balloons 4a and 4b. The catheter 2 is insertable towards a region of interest by "riding " on a pre-positioned guide-Aire 4c, This is a known procedure in interventional -cardiology, and therefore need not be elaborated further herein. To fix the relative position of the imaging probe 3 relative to the vesselwalls (not shown), the perfUsion balloons 4a and 4b are inflated until sufEcient contact is achieved between the balloons and the vessel walls. 7he positioning and operation of the imaging probe 3 permits imagihg of a region or slice 5 that intersects the vessel walls. The balloons positions are so designed that the shape distortion at boh distal and proximal planes of contact of y at the the wall wiJU have a negligible effect on the original wall morpholos; imaging sUce 5.

In. this specific example,- the -balloons 4a nd 4b have an annular shape and define openings 6a and 6b, respectively, and the imaging probe 3 is designed to have a lumen 6c so as to allow the continuous circulation of blood therethrougli, both during an image acquisition phase and at all other time periods. Rather, there are lumens through probe 3, Ahe imaging probe may include several irmer liquid contrasting agent, as weU passa,geways to pemixt passage of various liquids, e.g. a as throu,2b-flow of blood.

Further provi6d are leads 7 that nin along the intravascular catheter 2 to and from the imaging probe 3 to connect it to a motor drive unit 8 lodated at the catheter's rear portion. The motor drive urat 8 serves to rotate the - imaging probe 3 at the distal end of the catheter 2 around the Z-axis see- Fig, 2A), while image acquisition is in progress. lt should be noted, although not specifically showr4 that the motor drive unit 8 may also house- a transmit/receive switching unit a pre-amplificalion circuit or an amplifying unit for creating the required =ismission pulses. Alternatively, although not specifically showr the switching unit and the pre-amplifying components may be integrated. into the cattreter 2, in proximity to the imaging probe 3.

The motor drive unit 8 is controlled and powered by a spe-etrometer module 9 via cable 10. Receiving and transipitting channels, desigoated respectively Ila and llb are also connected to the spectrometer module 9, which controls the synchronized transmit pulse generation, signal acquisition, imaging probe rotation and, possibly -gradient activation and level, as the case may be. The _5 spectrometer 9 is Rirther connected to a control station 12 (e.g., a personal computer), through which the operator interfaces with the system.

The system I operates in the fbllow#ig-g manner, When triggered by the operator, the acquisition process takes place by running the imaging sequence and collecting received sivTWs while the imaging probe 3 is rotated, as will be described more specifically further below. In the end of the acquisition phase, the received signals are processed, and a bigh-resolution image of a cross section of t he vessel contained in the imaging slice 5 can be displayed on a monitor of the control station 12. Multiple cross section images can readily be collected by moving the imaging probe 3 along the Z-a3ds. It should be noteú1, allouggh not specifically shown, that the same can be achieved by using an array or "stack' of ets/RF coil units aligned along the Z-axis. The so-obtained cross section images can be -D images of the vessel.

brther Processed to reconstruct 3 It should be noted that the imagg method may not require the'rnechaffical rotation of the imaging probe 3. In this case.. the motor drive unit 8 can be replaced jo by an alternative catheter interface umi having -mostly electrical functions (not sho-om).

probe 3 is rhore specifically illustrated.

Turning now to Fig. 2A, the imaging The perfusion balloon s and the guide-wire, are not shown here, solely to simplify the illustration. of the main components of the Imaging probe 3. The probe 3 is located in a blood vessel lumen 13, and is surrounded by vessel walls 14. The minimally obstructed flow Of blood through the imaging probe 3 is sbo.xm by arrows F.

The imaging probe 3 comprises a magnetic field forming assembly composed of two permanent magnets 15 and 16 which, in the present example, have an annular shape, and are mounted on a hollow cylindrical magnetic core 17, leaving an ham-magnet gap 18 between them. The permanent nagnets, 15 and 16 can be made of a rare-earth magnetic material, such as NdFeB (Neodymium Ferrite Boron), and the magnetic core 17 can be made of a soft ferromagnetic material of relatively high permeability and high saturation flux d=ity such as sintered iron powder. Surroundbig inter-magnet: gap 18 is an annular virtual imaging slice 5.

It should however be noted that the permanen t mage gnets 15 and 16 and the magnetic core 17 may have non-circular crose,,-sections. "ib permanent magnets 15 and 16 may be fixed difFerently to the magnetic core 17, as well as be movable with respect to one another and to the magnetic core 17. Alternatively, the core may be located oE the Z-axis. may have a cross-section diflerent from that d s bed i c cri n the drawing, and the provision of any core is Optional, as Will be described below writh reference to Figs. 10a-10c. The permanent magnets may be solid (no lumen), have a different cross section or may be magnetized in. a digerent direction (for xample.

ridially outward/inward).

The Z-axis 23 is a symmetry axis of the cylindrical core 17 (and, conseguently, of each of the magnets 15 and 16), and is perpendicular to the X-axis 25 and the Y-ams 26. The permanent magnets. 15 and 16 are magnetized perpendicularly to the Z-axis, and in opposite directions to along the X-axis, i.e., each other: the N- and S-poles (not shown) of the magnet 15 face, respectively, the S- and N-poles (not shown) of the magnet 16.

As shown in Fig. 3, the magnetic configuration so created (i.e., by Te first rna?net 15, the second magnet 16 and -the. magnetic core 17)pr6duces a static metic field (i.e., a primary magnetic úeld). having two regions R, and R2 where m-a-tY stati magnetic.fild lines 22 are directed substantially. along the Z-axis 23. As 3 5 shown these reglons Ri and R2 are aligned along the X-axis. and located symmetrically with respect to the Z-axis, in the lateral space of the inter-magnet gap 18. The imaging, slice 5 is formed by regions where changes in the magnetic fields Z-component are substantially small, with respect to changes in. the

Z-Position. BY chanizing the size of the inter-magnet gap 18. the profile of the static magnetic field Bo in the imaging slice 5 can be varied. A cross section of the imaiping slice 5 in the figure plane is outlined.

Turning back- to Fig. 2, an.RF coil block 19, composed of an RF coil wound on a toroidal magnetic core, is mounted in the inter-mk2nets gap 18. As shown, the RF coil block 19 is designed so as to form a narrow coil. gap 19a, the purpose of which will be explained further below.

It should be noted, although not specifically shown, that an -gr adient coiT may be added to the RF coil blockr. Furthermore, am additional and. separate coil may be used for transmission. The transmission coil may be wound around the receiver coil, wound as opposed rings above and below the X-Y-plane, wound around the permanent magnets and core so to have a rectangular cross- section in the Y-Z-plane, etc, Some possible examples of the transmission coil geometry are described in the above prior art patents.

It should also be noted 1hat the device can be utilized for obtaining pseudo-one-dimensional cross sections of the blood vessel wall by positioning the RF-coil gap 19a in the requested radial direction (a set angle in the X-y- plane).

The received NMR signal can be used -to obtain one-dimensional. profiles of material density and other physical characteristics.

For a given imaging probe 3 having a 3mm outer diameter the static magnetic field Bo was calculated. at points along a radial vctor 25, which extends along the X-a)ds 25 starting at the X-position of the edges of the permanent graphically illustrated in Fig. 5. The magnets 15 and 16. The calculation results are calculated Z-component-of the static magnetic field B2; as a function of the distance along the -radial vector 25 (X-axis) can be seen. It is readily noticed that the mgmetic field strength decreases substantially with the increase of the radial distance from the Z-axis. creating a static magnetic field gradient aB:. According to the present invention, this magnetic field gradient is employed to advantage for imaggging, purposes, as will be described more specifically further below. It should be understood that the field profile depends on magnetic materials Used in the magnetic field forming assembly.

In order to excite the nuclei in the imaging, slice 5, an oscillating (RF) magnetic field, oriented perpendicularly to the static magnetic field, is, in one embodiment of the invention. created by the RF coil block 19 placed in the gap 18 between the magnets 15 and 16.

2B, the RF coil 19 is preferably wound around the toroidal -'5 Referring to Fig core 31, as illustrated. by a few representative coil windings 30. The toroidal core 31 is made of a ferrite material and has the narrow core gap 19a. The winding method can be either in a. single direction along the toroid perimetc3; 'in. one or more layers, in groups of winding, or in any other manner well known in the art. The coilducting wire can, alternatively, be wound in opposite directions, narnely a right-hand-helix from the gap 19a alonghalf of the toroid perimeter, and therefrom reversing winding direction- to a left-hand-helix up until. the other side of the gap 19a.

As shown in Fig. Ik. the RF-coil 19 is preferably fixed to the magnetic core 17 so that the coil. gap 19a is aligned with the X-axis 25. In other embodiments, such as where the static magnetic field is synimetric, around the Z-axis, the RF-coil

19 may rotate separately from the rest of the probe 3 components, Fig. 4 illustrates a cross section of a preferred embodiment of the RF coil unit 19 in the X-Y ('imaginf) plane, -the X and Y axes marked 2 and 26 respectively. According to this specific example, the RF coil unit 19 is. used for signal receptior4 and may also be used for winsmission of excitation pulses. As indicated above, the construction of 1he imaging, probe may be such that. a separate coil. is used for transmission. The magnetic core 17 is - not shown in the cross-section for clarity. The RF Coil unit 19, having a smooth circular cross-section in this embodiment, may alternatively have different cross-sections. The coiJ.

windings 30 may also vary in number geometxy. etc.

For illustrating the contribution of the RIF coil 19 having the above-describcd geometry to its sensitivity when used as a reception coil, a situation. in which a current is driven through the coil should be discussed. When used as a reception coD, a current is obviously not driven into the coU., but is rather induced in the coil by variations in extemal nuclear magnetization. However, the principle of reciprocity states that coil sensitivity to magnet ic field variations at a certain point in space is directly proportional to the magnetic field produced at that same point by a unit current driven through the coil. The discussion below should be understood as an illustration only for describing coil sensitivity wlien used for reception purposes. In embodiments where the same coil is used for twmnission, the situation of a current driveninto the coil can be taken literally.

When a current is driven through the coil windin, 30. a magnetic flux is gs created in the ferrite material through the toroidal core 31 and across the coil gap 19a. In a figurative way of speakin& the ferrite material in the toroidal core 31 e&ctively collects the ma-gaetic flt;x produced by the coil. current and concentrates it in the coil gap 19a. Stray -flux lines 32 in. the vicinity of the coil gap 19a produce a transverse magnetic field in the X-f plane of a relatively high strength in comparison with alternative coil designs of similar dimensions, such as solenoids or saddle coils. When used for reception of alternating magnetic fields, such cog geometry provides an improved coil sensitivity, which allows high- resolution images to be obtained even at highly inhomogencous magnetic fields. The segment-like region 33 is a-region in the imaging slice 5 where the magnetic field is

3 0 of sufficient strength,, or, wh= used for reception, where sensitivity is sufficient.

Varying the coil gap 19a size effects the RF magnetic field strength (or sensitivity) in the region 33.

It should be noted, although not.specifically shown, that in embodiments where the conducting wire is wound in the opposed manner, as described above, 1.5 the magnetic flux lines crate a different pattern, having a predominmt radial component along the X-axis.

Fig. 6 is a graphical representation of the RF magnetic field strength per -unit current as calculated for vanous points along the radial vector 25 (X- axis) -of an imaging probe 3 having a 3=. outer diameter. It can readily be seen that the RF magnetic field is highly non-hornogmeous. a factor which has to be taken into account in the imagang sequence and in a further signal processing method, as will be described Airther below. It should beunderstood that the present example relates to a specific choice of the geometry of the coil block 19, ferrite rnateri4 a coil winding method and the number of windings.

An exemplary imaging method that may be used in accordance with the present invention is based on a sequence illustrated in, Fig. 7, which is a variant of the Ca=-Purcell-Meiboom-Gill (CPMG) sequence. The terminology and graphical method used to describe this sequence is known per se, and therefore need not be specifically described. The first graph 61 describes the time-base envelope of the transmitted RF pulses used to excite theiauclei. The second graph 62 is a qualitative.

sketch of the expected nagnctic resonance signal. A series 63 of wideband, non-selective (or "hard") RF pulses is preferably used to simultaneously excite nuclei in the entire imaging sector 33, so as to obtain a series of spin- echoes 67.

One alternative for obtaining the spin-echoes 67 is the transmission of a 901IRP pulse 65 followed by a series of 180cRF pulses 66, using -the time delays saven M graph 61. It should, however, be noted that numerous other excitation schemes are applicable for obtaining a -series of spin-echoes 67, and can therefoie be used for the Purposes of the present invention. Sequences based on magmetizati on tilt angles different than 9011and 1900, on stimulated echoes, and on variations in pulse timing and phase are few examples of many dif[erent sequences well known in the art, which can be used in accordance with the present invention.

The spin-spreading, which is rephased at each echo center; stems predominantly from the strong x-gradient in the static mapetic field,, WMcb. exists permanently in the imaging sector 33. By reference to Fig. 5, whicb relates to one specific example of the present invention, this -gradient can be estimated at around I OOTesla/mcter, corresponding to a mapetic resonance ftequency range of 3.5NERz to 8.5NIHz for nuclei in an imaging sector 33 of a 2mm radial dimension.

It is thus evident that the methcO according to the invention enables to simultaneously operate within a substantially wide frequency ban&-,idth of the magnetic resonance signals with respect to the m= &cquency value. In the present example, this fTequency bandwidth is approximately 100% of the mean frequency value. In general, the method of the invention enables to successfully operate in most any non-homogmeous magnetic field that ensures a sufficiently high signal-to-noise ratio, whereas conventionally know-a techniques are limited to the NMR frequency bandwidth of less than 1% of the mean frequency valuc, It will be understood that one-dimensional (1-D) spin-density profiles of the imaging sector 33 can readily be obtained by sampling the echo signal (67a in Fig.

7, for example), using typical r values and acquisition window sizes of several microseconds, and transfmm-Ang it using a Fourier transform based aleorithm. The I fact that the static magnetic field profile (Fig. 5) is not perfectly Unew can be readily compensated for duringpost-processing, The re latively short c values make possible the acquisition of a few thousands of spin-echoes 67 in one excitation series 63, having the duration of a typical transverse relaxation time (known as T2). The multiplicity of spin-e choes 67 acquired. can be averaged prior to any tansformation in order to improve the signal-to-noise (S/N) ratio and, as a result, the quality of the final image.

In accordance with the -present invention; scveral methods are available for obtaining two dimensional (2-D) images of the imagging slice 5. One type of methods utilizes an integrated -gradient coil, of whicb me suggested design is disclosed in the above-indicated US Patent No. 5,572,132. Simpler gradient coil designi are also possible since it is sufficient to control the gradient profile over the imaging sector 33 alone. An. alternative method for obtaining lateral () separation-, which may obviate the need for a -gradient coil, utilizes processing of phase variations in the receive signals acquired from overlapping imaging sectors (not shown) as the imagiag- probe 3 is rotated around its axis 23. Possible V pulse series to be used for exciting nuclei in consecutive, preferably overlapping, irnaging sectors are represented by 63 and 64 in Fig. 7. 7be time separation between the series 63 and 64 is labeled TR.

In both methods, a slow rotation of the imaging probe can be used to 001lect data from imaging sectors 33 covering the entire imaging slice 5. This data ran then be processed and displayed as a 2-D image.

Two aspects of the previously mentioned RF magnetic field inhomogeneity

AU now be considereA Me first one covers the effect of non-uniform excitation in -'5 wide-band excitation schemes, for cases where the RF coil block 19 having the filed profile of Fig. 6 is utilized for t=smission purposes. Several phase-cycling techniques exist and are well 'known in the art, which can effectively compensate for non-uniform excitation. An alternative. method may be emp loyed for mmimizmg the eff=five RF magnetic field inhomogcneity by using tailored -3o trammissionRF wavefbrrns.,InrefercncetoFigs. 5 and 6, aspect-al distribution of the RF magnetic field strength can be calculated Le-the relative RF magnetic field strength at each magnetic resonance frequency. Inverting the calculated spectral distribution yields the pulse spectrum required for obtain lg a uniform excitation.

Fig. 8.Wustrates a _graph 70 corresponding to an RF waveform calculated in s this manner'for. the typical magnetic field profiles given by the graphs of Figs. 5 and 6. The calculated RF (current) waveform is displayed in relative units vs. a time base in rnicroseconds. Replacing the rectangular enveloped RF pulses 65 and 66 in the sequence 60 by the calculated RF waveform of an appropriate level will create an effiectively uniform RF magnetic field throughout the,c,, sector 33.

Fig. 9 fflustrates the relative R.F magnetic field strength 82, created by the

1 calculated RF waveform vs. the magnetic resonance frequency scale 81.

Uniformity of the rel ative RF magnetic field strength 82 can be readily observed in the mapetic resonance frequency range of 3.5 to 8.5 MEIz, characterizing nuclei in the imaging sector 33 of an imaging probe 3 of 3n= outer diameter.

A second effect of RF magnetic field non-b.omogencity,.when the IF coil block is used for signal reception, is the existence of spacially dependent coil sensitivity. The effect of variable sensitivity on signal intensity across the image can be readily compensated for durinc,,, the processing stage, at the expense of enhanced noise levels in image regions characterized bypoor sensiti-vity.

Turning now to Fig. 10a-10c, there are illustrated three dillerent examples, respectively, of alternative constructions of an. imaging probe suitable to be used in the MRI-based system, for example. the system of Fig. 1.

According to the example of Fig. 1 Oa, an imaging probe 103 comprises two permanent magnets 115 and 116 mounted.. on a magnetic core 117, which, in distinction to the previously described example, connects the periphery regions of the mamets rather than their central regions. The magnets are magnetized along the X-axis, the magnet 115 along the +X and the magnet 116 - along the -X direction, and spaced from each other along the Z-axis to form an inter-magnet gap. 118 in which an RF coil block 119 is accommodated. As shown, the coil block.119 has a coil gap 119a.

In the example of Fig. 10b, an imaging probe 203 comprises a magaetic field forming assembly formed by a single permanent magnet 215 having a thinner region 218 between its two- poles 215a and 215b. The magnet 215 is magrietized along the Z-axis- An RF coil block 219 is positioned in the X-Y plane such that its coil gap 219a is located proximate to the magnet gap 218.

According to the example of Fig. I Oc, an imaging probe 303 comprises two spaced-apart axially oriented permanent magnets 315 and 316 mounted on a magnetic core 31.7, for example, connecting their central regions. As for the RF coil block 319, it is formed with four spaced-apart coil gaps 319a, 319b.. 319c and j.o 319d.

It should be understood that the above constiuction of the RF coil block 319 is suitable for any possible geometry of the magnetic field forming assembly,

Moreover, the U coil block according to the invention, namely formed with at least one coil gap arranged in a plane substantially perpendicular to the plane (Z-plane) can be advantageously used with any magnetic field forming assembly produLga either homogeneous or non-homogreneous primary magnetic field.

Generally speaking, any design of the probe is possible, provi ded ffiat. the RF coil block has the region of the sufficiently higb sensitivity, the primary static magnetic field is substantially perpendicular to the RF field (which would have been created if a current was to be driven into the RF coil) at any point in this rqzioi4 and the combination of the static field intensity and the RF coil sensitivity is sufficiently high to obtain the required signal-to-noise ratio. For imaging purposes, the static field patt= (combined with field gradients, if any) should allow the i.e., points which are to be resolved in the image can be required mapping subjected to distinct static field values.

In accordance xith the present invention, the MRI probe enables imaging either from within cavities in the human body, or by immersion in any medium to be tested, or in any other industrial application. Adaptation of the invention for a specific application can bc accomplished by means of variations in most of the imaging probe characteristics. Also, the imaging method can be varied in accordance with the imaging requirements of the specific application.

The advantages of the present invention. are thus self-evident. The provision of a sufficiently high sensitivity receiving coil block enables operation with an extremely non-homogeneous static magnetic field. The non-homogeneity of the static magnetic field allows for advantageous use of more extremely multiple echo averaging for imaging purposes. The provision the high sensitivity reception. coil enables to significantly improve the image resolution. Pulse shaping al. lows for compensating for nor)-homogerLeous t-ansmission RF fields. The single-gAp coil block enables to obtain I-D images without the coil rotation, and rotation is required in cases wh ere 2-D imaging is to be achieved. Rotation is required only for those magnetic field sources, which create fiel-& asymmetric for rotation about the

Z-axis. If the static field is symmetric, the permanent magnet(s) may not be rotated.

Similarly, if multiple coil gaps create imaging sectors which cover the entire 3600 space, the coil Iock may not be rotated. Rotation of the permanent magnets and/or the coil block is therefore only a mmor and certainly not an essential part of the invention.

Those skilled in the art will readily appreciate that various modifications and changes can be applied. to the prefer:red embodiments of the invention as hereiribefore exempMed without departing from its scope defined in and by the appended claims.

Claims

CLAIMS:

1. A mefflod for detectingg sigmjs coming from a medium, the method comprising:

(i) producing. a primary, substantially non-homogeneous, c%-temal magnetic field in the medium; (U) detecting the magnetic resonance signals fl-6i within at least one region of said primarn substantially non-homog-encous magnetic field.

2.. The method according to Claim 1, wherein the magnetic resonance signals.

are simultaneously detected from said at least one region.

3. Tle metliod according to Claim 2, wherein said simultaneously detected NMR signals = of a substantially wide frequency bandwidth with respect to their mean frequency value.

4. The method according to Claim 3, Wherein said substantially wide tequency. bandwidth is greater than 1% of the mean frequency value.

1:

5 5. A probe for producing Nuclear Maffinetic Resonance) signals comikg from a medium surrounding the probe and detecting the produced signals, the probe comprising:

a magnetic Beld-formnag assembly that produces a primary, substantially non-homogeneous, external masnetic field; d an a transceiver unit cOmPrising at least one coil block capable of detecting maP'-llc resonance signals witilin at least one region of said primary magnetic field, said at least one region extending tom the probe up to distances substantially of the probe dimensions.

6. The probe according to Claim 5, wherein said at least one coil block is formed with at least one coil gaP, the gap plane being aligned. substantially parallel to the direction of the primary magnetic field in said at least one region.

7. The probe according to Claim 6, wherein said at least one region is in the form Of a segment extending radially from said at lcut one -coil gap.

S. Tile probe according to Claim 5. wherein said magnetic field forming assembly comprises two magnets mounted on.a common ' magnetic core and arranged. in a spaced-apft axial relationship along.. the Z-axis, defining an inter-magnet gap therebetween, said at least one segment extending. radially:Erom said at least one coil gap towards a slice defined by a radial space surrounding the inter-magnet gap

9. The probe according to 'Claim 8, wherein said magnets are magne&,ed in opposite directions along an axis perpendicular to the Z-axis, thereby producing the primary exwmal magnetic ' field substantially parallel to the Z-axis.

'

10. The probe according to Claim 8, wherein said magnets are radially magnetized in opposite directions, such that one of the magnets is magnetized in a radially inward dir6ction and the other magnet is magnetlZCd in the radially outwards direction.

11.. The probe according to Claim 5, being rotatable so as to provide rotation of said at least one region through the =ounding medium.

12.. The probe according to Claim 5, being dis1aceable relative to the surrounding medium.

13. The probe according to Claim 5, wherein said tmnscerver unit comprises at least one additional coil block capable of detecting the N-MR signals within at least :zo one addition repon of said primary magnetic field.

14. The probe according to Claim 5, and also comprising at least one additional magnetic field forming assembly with a corresponding at least one additional cog block.

15. The probe according to Claim 5, whereh). said at least one coil block. also serves for signal transmission,, being thus capable of generating an oscillating magnetic field producing said magnetic resonance signals.

16. The probe according to Claim 6, wherein said at least one coil block comprises a coil wound on a magnetic core ha---ving a non-continuous substantially toroidal shape defining said at least one coil W.

17. The probe according to Claim 8, wherein said magnets are substantially cylindricaL

18. Ile probe according to Claim 1T, wherein said magnets are tubular.

19. The, probe according to Claim 8, wherein said magenets are permanent magnets.

1

20. The probe according to Claim 5, for use in medical equipment for NMR signal detection. wherein said probe is a'catheter, said magnetic field forming assembly and said at least one RF coil block being housed within said cathetet

21. The probe according to Claim 20, having a hollow lumen to allow blood, flow,therelliough.

22. 'lIe probe according to Claim' 20, being integrated in a medical catheter-based therapeutic or diagiostic device.

23. A device for magnetic resonanc c imaging of a medium, the device comprising an imaging probe to be locate-d in the vicinity of said medium and 0, connected to a control station for generating transmission pulses, and for receiving, processing' and displaying data generated by the probe, the probe comprising:

(c) a magnetic field forming assembly that produces a primary, substantially non homogeneous, external magnetic field in the medium; and (d) a transceiver unit comprising at least one coil capable of detectiucc, magnetic resonance signals from within at least one region of said primary substantially non-homogeneous manetic field. said at least one region extending from. the probe up to distances substantially of the probe dimensions.

24. A device for NMR measurements in a medium, the device comprising a probe to be located in the vicinity of said medium and connected to a control station for generating transmission pulses, and for receiving, processing, and displaying data generated by the probe, the probe comprising..

(a) a magnetic field forming assembly that produces a primary, substantially non homogeneous, ex magnetic. field in the medium; and (b) a transceiver unit comprising at least one coil capable of detecting W magnetic resonance signals from within at least one region of said primary, substantially non-homogeneous rdagnetic:Geld, said at least one region. extending from the probe up to distances substantially of the probe dimensions.

25. A device for magnetic resonance unaging QvUU) and NMR measurements of a medium surrounding an imaging probe of the said device, the imaging probe comprising:.

a magaletic field-forming assembly defining a longitudinal, Z-axis, and comprising two magnets spaced-apart along the Z-axis to define an inter-magnet gap and an imaging slice within anannular radial spacearound the inter-magnetgap, wherein the magnet are magnetized so as to produce a primary substantially non-homogeneous external magnetic field; at least one coil block capable of inducing an oscillating magnetic field substantially orthogonal with respect to the Z-axis and capable of detecting magnetic resonance sign& within at least one region of said primafy magnetic field.. wherein the coil block is'accommodated in said inter- magnet gap and has at least one coil gap located -in a plane substantially cr perpendicular to the Z-axis such that the coil gap plane is substantially parallel to the Z-axis, said at least one region being located within a segment like region extending fi-om said at least one coil gap towards said imaging slice.

26.

A transceiver unit for use in a probe for detecting NMR signals of a surrounding medium- the transceiver unit comprising at least one coil block capable of detecting magnetic resonance signals within at I=t one region of a primary extemal magnetic field, wherein said coil block comprises a coil wound on. a substantially toroidal core having at least one core gap.

27. The unit according to Claim 26, wherein said at least one coil block induces A an oscillating magnetic field directed substantially orthogonal to the static primary magnetic faeld.